Electrochemical test sensor

ABSTRACT

The present invention relates to electrochemical sensor strips and methods of determining the concentration of an analyte in a sample or improving the performance of a concentration determination. The electrochemical sensor strips may include at most 8 μg/mm2 of a mediator. The strips, the strip reagent layer, or the methods may provide for the determination of a concentration value having at least one of a stability bias of less than ±10% after storage at 50° C. for 4 weeks when compared to a comparison strip stored at −20° C. for 4 weeks, a hematocrit bias of less than ±10% for whole blood samples including from 20 to 60% hematocrit, and an intercept to slope ratio of at most 20 mg/dL. A method of increasing the performance of a quantitative analyte determination also is provided.

REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.14/964,721 filed Dec. 10, 2015, which has been allowed; U.S. patentapplication Ser. No. 14/964,721 is a continuation of U.S. patentapplication Ser. No. 14/252,226 filed Nov. 14, 2013 and issued as U.S.Pat. No. 9,239,312; U.S. patent application Ser. No. 14/252,226 is adivisional of U.S. application Ser. No. 14/079,922 filed Nov. 14, 2013and issued as U.S. Pat. No. 8,728,299; U.S. application Ser. No.14/079,922 is a divisional of U.S. application Ser. No. 12/951,382 filedNov. 22, 2010 and issued as U.S. Pat. No. 8,702,965; U.S. applicationSer. No. 12/951,382 is a divisional of U.S. application Ser. No.11/853,010 filed Sep. 10, 2007 and issued as U.S. Pat. No. 7,862,696,which claims the benefit of U.S. Provisional Application No. 60/846,688entitled “Biosensor System Having Enhanced Stability and HematocritPerformance” filed Sep. 22, 2006, all of which are incorporated byreference in their entireties.

BACKGROUND

Biosensors provide an analysis of a biological fluid, such as wholeblood, urine, or saliva. Typically, a biosensor analyzes a sample of thebiological fluid to determine the concentration of one or more analytes,such as glucose, uric acid, lactate, cholesterol, or bilirubin, in thebiological fluid. The analysis is useful in the diagnosis and treatmentof physiological abnormalities. For example, a diabetic individual mayuse a biosensor to determine the glucose level in whole blood foradjustments to diet and/or medication.

Biosensors may be implemented using bench-top, portable, and likedevices. The portable devices may be hand-held. Biosensors may bedesigned to analyze one or more analytes and may use different volumesof biological fluids. Some biosensors may analyze a single drop of wholeblood, such as from 0.25-15 microliters (μL) in volume. Examples ofportable measurement devices include the Ascensia Breeze® and Elite®meters of Bayer Corporation; the Precision® biosensors available fromAbbott in Abbott Park, Ill.; Accucheck® biosensors available from Rochein Indianapolis, Ind.; and OneTouch Ultra® biosensors available fromLifescan in Milpitas, Calif. Examples of bench-top measurement devicesinclude the BAS 100B Analyzer available from BAS Instruments in WestLafayette, Ind.; the CH Instruments' Electrochemical Workstationavailable from CH Instruments in Austin, Tex.; the CypressElectrochemical Workstation available from Cypress Systems in Lawrence,Kans.; and the EG&G Electrochemical Instrument available from PrincetonResearch Instruments in Princeton, N.J.

Biosensors usually measure an electrical signal to determine the analyteconcentration in a sample of the biological fluid. The analyte typicallyundergoes an oxidation/reduction or redox reaction when an input signalis applied to the sample. An enzyme or similar species may be added tothe sample to enhance the redox reaction. The input signal usually is anelectrical signal, such as a current or potential. The redox reactiongenerates an output signal in response to the input signal. The outputsignal usually is an electrical signal, such as a current or potential,which may be measured and correlated with the concentration of theanalyte in the biological fluid.

Many biosensors have a measurement device and a sensor strip. A sampleof the biological fluid is introduced into a sample chamber in thesensor strip. The sensor strip is placed in the measurement device foranalysis. The measurement device usually has electrical contacts thatconnect with electrical conductors in the sensor strip. The electricalconductors typically connect to working, counter, and/or otherelectrodes that extend into a sample chamber. The measurement deviceapplies the input signal through the electrical contacts to theelectrical conductors in the sensor strip. The electrical conductorsconvey the input signal through the electrodes into a sample depositedin the sample chamber. The redox reaction of the analyte generates anoutput signal in response to the input signal. The measurement devicedetermines the analyte concentration in response to the output signal.

The sensor strip may include reagents that react with the analyte in thesample of biological fluid. The reagents may include an ionizing agentfor facilitating the redox reaction of the analyte, as well as mediatorsor other substances that assist in transferring electrons between theanalyte and the conductor. The ionizing agent may be an oxidoreductase,such as an analyte specific enzyme, which catalyzes the oxidation ofglucose in a whole blood sample. The reagents may include a binder thatholds the enzyme and mediator together.

One disadvantage of the reagent compositions used in conventionalbiosensors is the change in measurement performance, either accuracy orprecision, that occurs when the sensor strip is stored. The electronicsand analysis methods used by the measurement device to determine theanalyte concentration of the sample are generally selected in view ofthe reagent composition on the sensor strip performing as initiallymanufactured. However, after transportation and storage on storeshelves, the reagent composition degrades with time and temperature.This change in the chemistry of the reagent composition may result in areduction of measurement performance.

To increase the long-term stability of biosensor reagent compositions,conventional biosensors generally rely on a substantial excess of enzymeand mediator in relation to the amount of these reagents required toanalyze the sample. Expecting these reagents to degrade over time,conventional reagent compositions include substantially greater amountsof enzyme and/or mediator than required to stoichiometrically react withthe analyte. In addition to increasing the cost of the biosensor throughthe use of sacrificial reagents, the unnecessary reagents may require alarger sample volume, longer analysis time, and decrease the measurementperformance of the biosensor due to many factors.

For example, PCT publication WO 88/03270 discloses an overall depositiondensity of 3 mg/cm² (30 μg/mm²) with a screen printing method. Therelative amount of K₃Fe(CN)₆ was 57.7%, phosphate buffer at 28.8%, andglucose oxidase (GO) at 3.6%. Translating these percentages intodeposition densities on the sensor strip results in a K₃Fe(CN)₆ densityof 17.31 μg/mm², a phosphate buffer density of 8.64 μg/mm², and a GOdensity of 1.08 μg/mm². In another example, column 17, lines 25-35 ofU.S. Pat. No. 4,711,245 discloses the deposition of 15 μL of a 0.1 Msolution of 1,1′-dimethylferrocene in toluene onto a disk electrodehaving a diameter of 4 mm. With a molecular weight of 214 M.U., the1,1′-dimethylferrocene mediator was applied at a deposition density of25.5 μg/mm² [(15 μL*0.1M*214 g/mol)/2²*3.14 mm²=25.5 μg/mm²]. In afurther example, U.S. Pat. No. 5,958,199 discloses the deposition ontothe sensor electrode of 4 μL of a solution including 40 mg of GO, 16 mgof K₃Fe(CN)₆, and 20 mg of CMC in 1 mL of water. In this instance, thedeposition densities were 6.67 μg/mm² for GO, 10.67 μg/mm² forK₃Fe(CN)₆, and 13.33 μg/mm² for CMC with an estimated electrode area(deposition area) of 6 mm². In a further example, U.S. Pat. No.5,997,817 describes a reagent formulation including 59 g of K₃Fe(CN)₆dissolved in approximately 900 mL water with other ingredients.Approximately 4.5 μL of this reagent was deposited onto a 21.4 mm²(3.2×6.7) opening to give a mediator deposition density of 13.96 μg/mm²(4.5×10⁻³ mL*59 g/900 mL).

In each of these examples, the reagent compositions had mediatordeposition densities in the 10-25 μg/mm² range, while the enzymedeposition density was in the 1-6 μg/mm² range. This large mediatorloading in relation to the enzyme may be attributable to the singleapplication of the composition to both the working and counterelectrodes. Depending on sensor design, mediator may function at thecounter electrode to support the electrochemical activity at the workingelectrode. Thus, a single reagent composition deposition covering bothelectrodes may result in substantially overloading the working electrodewith mediator.

The examples show that excesses of enzyme and mediator are used toensure that enough active ingredients are present for accurate glucosemeasurement. Using sensor strips manufactured with increased sacrificialamounts of reagents after long-term storage may result in thedisadvantage of a drift in measurement performance. This drift may beobserved in at least two ways: (1) a background current increase overtime (affecting the calibration intercept) and (2) a shift in sensorsensitivity (affecting the calibration slope).

During storage, reduced mediator may be produced from interactionsbetween the oxidized mediator and the enzyme system and polymer. This isa natural process believed to be governed by thermodynamics. The largerthe amount of mediator or enzyme, the larger the amount of reducedmediator that is produced. As the concentration of reduced mediatorincreases over time, the background current will increase toward the endof the shelf-life of the sensor strips.

Multiple methods have been proposed to reduce the effect of drift onsensor performance before use of a stored sensor strip. For example,Genshaw et al. in U.S. Pat. No. 5,653,863 disclosed a method of using arelatively long initial pulse before the analysis to oxidize mediatorthat was reduced during transport and storage. While effective, thismethod lengthened the time required to complete the analysis.

Thus, it would be desirable to increase the long-term stability of thereagent composition to improve the measurement performance of thebiosensor after transportation and storage. Such a long-term stabilityincrease of the reagent composition may increase the measurementperformance of the biosensor and provide a longer shelf-life for thesensor strips. It also would be desirable to reduce the amount ofsacrificial enzyme and/or mediator included in the reagent compositionand to decrease the time required to complete the analysis.

Another drawback of conventional biosensors used to measure the glucoseconcentration in whole blood (WB) samples is referred to as the“hematocrit effect.” In addition to water and glucose, WB samplescontain red blood cells (RBC). Hematocrit is the volume of a WB sampleoccupied by RBC in relation to the total volume of the WB sample and isoften expressed as a percentage. The hematocrit effect occurs when redblood cells block the diffusion of the analyte and/or mediator to one ormore electrodes of the biosensor. Since the output signal measured bythe biosensor corresponds to the rate of diffusion of the analyte and/ormediator, the RBC may introduce error to the analysis by interferingwith this diffusion process. Thus, the greater the hematocrit percent(volume of red blood cells) deviates from the %-hematocrit systemcalibration for a WB sample, the greater the hematocrit bias (error) inthe glucose readings obtained from the biosensor.

WB samples generally have hematocrit percentages ranging from 20 to 60%,with ˜40% being the average. If WB samples containing identical glucoselevels, but having hematocrits of 20, 40, and 60%, are tested, threedifferent glucose readings will be reported by a system based on one setof calibration constants (slope and intercept of the 40% hematocritcontaining WB sample, for instance). Even though the glucoseconcentrations are the same, the system will report that the 20%hematocrit WB sample contains more glucose than the 40% hematocrit WBsample, and that the 60% hematocrit WB sample contains less glucose thanthe 40% hematocrit WB sample due to the RBC interfering with diffusionof the analyte and/or mediator to the electrode surface. Thus,conventional biosensors may not be able to distinguish between a loweranalyte concentration and a higher analyte concentration where the RBCinterfere with diffusion.

Conventional biosensors are generally configured to report glucoseconcentrations assuming a 40% hematocrit content for the WB sample,regardless of the actual hematocrit content. For these systems, anyglucose measurement performed on a blood sample containing less or morethan 40% hematocrit will include some hematocrit bias attributable tothe hematocrit effect.

Various methods and techniques have been proposed to reduce the bias ofthe hematocrit effect on glucose measurements. For example, Ohara et al.in U.S. Pat. No. 6,475,372 disclosed a method of using the ratio ofcurrents from a forward and a reverse potential pulse to compensate forthe hematocrit effect. McAleer et al. in U.S. Pat. Nos. 5,708,247 and5,951,836 disclosed a reagent formulation using silica particles tofilter the RBC from the electrode surface for reducing the hematocriteffect. Carter et al. in U.S. Pat. No. 5,628,890 disclosed a method ofusing wide electrode spacing in combination with mesh layers todistribute the blood sample to reduce the hematocrit effect.

These conventional techniques for reducing the bias attributable to thehematocrit effect included (a) co-deposition of a polymer to minimizethe hematocrit effect, (b) addition of various kinds of fused silica toenhance the filtration effect for the polymer layer, (c) compensationcoefficients based on the ratio of currents from a forward and a reversepotential pulse, and (d) self-compensation by utilizing the existingsolution resistance of the whole blood samples. Although these methodsmay be useful, conventional glucose sensors continue to exhibitsignificant analytical bias attributable to the hematocrit effect,generally from about 15 to 30%. Thus, it would be desirable to providesystems for quantifying analytes in biological fluids, in particular theglucose content of whole blood, which reduces bias from the hematocriteffect.

SUMMARY

In one aspect, an electrochemical sensor strip includes a base, firstand second electrodes on the base, and a lid on the base. The stripincludes at least one first layer on a first conductor, the first layerincluding a reagent layer including at most 8 μg/mm² of a mediator. Thestrip provides a determined concentration value having at least one of astability bias of less than ±10% after storage at 50° C. for 2 weekswhen compared to a comparison strip stored at −20° C. for 2 weeks, ahematocrit bias of less than ±10% for whole blood samples including from20 to 60% hematocrit, and an intercept to slope ratio of at most 20mg/dL.

The first and second electrodes of the strip may be in substantially thesame plane and the second electrode may include the first layer on asecond conductor. The second electrode may include a second layer on thesecond conductor, and the second layer may include a reagent layerdifferent in composition from the reagent layer of the first layer. Theelectrodes may be separated by greater than 200 μm and may be separatedfrom an upper portion of the lid by at least 100 μm.

The average initial thickness of the reagent layer of the strip may beless than 8 μm or may be from 0.25 to 3 μm. The reagent layer of thestrip may be formed at a deposition density of at most 0.2 μL/mm² from areagent solution. The reagent layer may include poly(ethylene oxide),polyvinyl alcohol, hydroxyethylene cellulose, carboxy methyl cellulose,or a combination thereof as a polymeric binder. The deposition densityof the polymeric binder may be at most 2 μg/mm² on the first conductor.The polymeric binder may be partially water soluble and/or may form agel-like material on hydration.

The reagent layer may include an enzyme system at a deposition densityof at most 0.8 μg/mm² and/or include at most 1.3 Units of enzyme. Thereagent layer on the working electrode may include at most 2/g/mm² ofthe mediator. The mediator may be a two electron transfer mediator andmay be 3-phenylimino-3H-phenothiazines, 3-phenylimino-3H-phenoxazines,salts thereof, acids thereof, derivatives thereof, or combinationsthereof. The mediator may have a redox potential at least 100 mV lowerthan that of ferricyanide.

The strip may have a stability bias less than ±5% after storage at 50°C. for 2 or 4 weeks when compared a comparison strip stored at −20° C.for 2 or 4 weeks, respectively. The strip may have a hematocrit biasless than +5% for whole blood samples including from 20 to 60%hematocrit. The strip may have an intercept to slope ratio of at most 10mg/dL or at most 1 mg/d L.

In another aspect, an electrochemical sensor strip includes a base,first and second electrodes on the base, and a lid on the base. Thestrip includes at least one first layer on a first conductor, the firstlayer including a reagent layer including a mediator and an enzymesystem, where the reagent layer provides a determined concentrationvalue having at least one of a stability bias of less than ±10% afterstorage at 50° C. for 2 weeks when compared to a comparison strip storedat −20° C. for 2 weeks, a hematocrit bias of less than ±10% for wholeblood samples including from 20 to 60% hematocrit, and an intercept toslope ratio of at most 20 mg/dL.

In another aspect, a method of determining the concentration of ananalyte in a sample is provided. The method includes applying a pulsesequence to the sample, the pulse sequence including at least 3 dutycycles within 30 seconds. The method also includes determining theconcentration of the analyte in the sample, where the concentration hasat least one of a stability bias of less than ±10%, a hematocrit bias ofless than ±10% for whole blood samples over a 20 to 60% hematocritrange, and an intercept to slope ratio of at most 20 mg/dL.

The method may include at least 3 duty cycles within 9 seconds and thepulse sequence may be complete in at most 5 seconds. Determining theconcentration of the analyte in the sample may include determining theconcentration of the analyte in the sample from a current measurementtaken within 2 seconds from the application of the pulse sequence. Eachduty cycle may include an excitation and a relaxation, where eachexcitation may have a duration from 0.01 to 3 seconds. The excitationsmay have a summed duration of at most 10 seconds or at most 2 seconds,and the excitations may have amplitudes differing by 500 mV. Theexcitations may be at most 45% of the time of the pulse sequence. Therelaxations each may have a duration of at least 0.2 seconds or may havea duration of from 0.2 to 3 seconds. The pulse sequence may include aninitial excitation from 0.75 to 3 seconds in duration, where thisinitial excitation is longer in duration than the excitations of theduty cycles.

The method may determine a concentration having a stability bias lessthan ±5% after storage at 50° C. for 2 or 4 weeks when compared acomparison strip stored at −20° C. for 2 or 4 weeks, respectively. Theconcentration may have a hematocrit bias less than ±5% for whole bloodsamples including from 20 to 60% hematocrit. The concentration may havean intercept to slope ratio of at most 10 mg/dL or at most 1 mg/dL.

In another aspect, a method of increasing the performance ofquantitative analyte determination includes introducing an analytecontaining sample having a liquid component to an electrochemical sensorstrip, the strip having a base, a first conductor on the base, a secondconductor on the base, and at least one first layer on at least thefirst conductor, where the at least one first layer includes a reagentlayer including a polymeric binder and the sample provides electricalcommunication between the first and second conductors. The method alsoincludes applying an electric potential between the first and secondconductors in the form of at least 4 read pulses within 30 or within 9seconds and measuring at least one of the read pulses to provide aquantitative value of the analyte concentration in the sample withincreased performance attributable to at least one performance parameterselected from a stability bias of less than ±10% after the strip isstored at 50° C. for 2 weeks when compared to a comparison strip storedat −20° C. for 2 weeks, a hematocrit bias of less than ±10% over a 20 to60% hematocrit range for whole blood samples, an intercept to sloperatio of at most 10 mg/dL, and combinations thereof.

The read pulses may be complete in at most 5 seconds, the read pulsesmay be measured within 2 seconds of applying the electric potential, andthe read pulses may each have a duration of from 0.01 to 3 seconds. Theread pulses may have a duration of at most 2 seconds and may be ofamplitudes having a difference within 500 mV.

The method may determine the quantitative value having a stability biasless than ±5% after storage at 50° C. for 2 or 4 weeks when compared acomparison strip stored at −20° C. for 2 or 4 weeks, respectively. Themethod may determine the quantitative value having a hematocrit biasless than ±5% for whole blood samples including from 20 to 60%hematocrit. The method may determine the quantitative value having anintercept to slope ratio of at most 10 mg/dL or at most 1 mg/dL.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention can be better understood with reference to the followingdrawings and description. The components in the figures are notnecessarily to scale, emphasis instead being placed upon illustratingthe principles of the invention. Moreover, in the figures, likereferences numerals generally designate corresponding parts throughoutthe different views.

FIG. 1A is a perspective representation of an assembled sensor strip.

FIG. 1B is a top-view diagram of a sensor strip, with the lid removed.

FIG. 2A is an end view diagram of the sensor strip of FIG. 1B.

FIG. 2B represents the transfer of a single electron by a mediator froman enzyme system to a working electrode.

FIG. 2C represents the transfer of two electrons by a mediator from anenzyme system to a working electrode.

FIG. 3 represents an electrochemical method of determining the presenceand concentration of an analyte in a sample.

FIGS. 4A-4D depict examples of gated amperometric pulse sequences wheremultiple duty cycles were applied to the sensor strip after introductionof the sample.

FIGS. 5A and 5B depict the dose response curves for sensor stripsincluding the PQQ-GDH enzyme system at the 20%, 40% and 55% hematocritlevel.

FIG. 5C presents I/S ratios in mg/dL determined from experimental data.

FIG. 5D shows I/S ratios from 0 to 20 mg/dL for multiple glucoseconcentrations determined from sensor strips having reagent compositionsRC2, RC3, or RC4.

FIG. 5E shows the nearly identical hematocrit performance of the systemwith plasma and 40% hematocrit whole blood samples.

FIG. 6A shows the stability bias for four reagent compositions includingthe PQQ-GDH enzyme system after 2 weeks at 50° C.

FIG. 6B shows the stability bias for four reagent compositions includingthe PQQ-GDH enzyme system after 4 weeks at 50° C.

FIG. 6C shows the stability bias for three reagent compositionsincluding the PQQ-GDH enzyme system after 52 weeks at 25° C. under 80%relative humidity.

FIG. 6D shows the variance in stability bias for sensor strips includingRC2 where the data point used to calculate the analyte concentration wastaken at varying times after starting the analysis.

FIG. 7A depicts the dose response curve from a sensor strip for multiplewhole blood samples.

FIG. 7B depicts the absolute hematocrit bias spreads for fourmanufacturing lots of sensor strips obtained with multiple whole bloodsamples.

FIG. 7C compares the hematocrit sensitivity between present andconventional sensor strips.

DETAILED DESCRIPTION

Biosensors provide patients with the benefit of nearly instantaneousmeasurement of glucose levels. Errors in these measurements may beattributable to a degradation of the reagent composition and/or thehematocrit effect. Degradation of the reagent composition is acontinuous process that occurs during the time period that the sensorstrip is transported and stored after manufacture. Multiple factors mayaffect the rate at which the reagent composition degrades, includingtemperature. The hematocrit effect arises when red blood cells randomlyaffect the diffusion rate of measurable species to the conductor surfaceof the working electrode.

By reducing the amount of mediator and/or enzyme used on the sensorstrip, the long-term stability of the reagent composition may beincreased in relation to conventional biosensors and reagentcompositions. Thus, the stability bias and intercept to slope ratios ofthe sensor strip may be improved. Furthermore, by combining a gatedanalysis method with the stability-enhanced reagent composition, thehematocrit effect may be reduced. Thus, one or any combination of theseand other performance parameters may be improved in accord with thepresent invention.

In one aspect, the biosensors of the present invention demonstrate astability bias of preferably less than ±10%, more preferably less than±5%, after storage at 50° C. for 4 weeks when compared to sensor stripsstored at −20° C. for 4 weeks. In another aspect, the biosensors of thepresent invention demonstrate a hematocrit bias of preferably less than±10%, more preferably less than ±5% for WB samples including from 20 to60% hematocrit. In another aspect, the biosensors of the presentinvention preferably demonstrate an intercept to slope ratio of at most20 mg/dL, more preferably at most 10 mg/dL or at most 6 mg/dL, and evenmore preferably at most 1 mg/dL. These and other performance parametersof the sensor strip may be improved.

FIGS. 1A and 1B depict a sensor strip 100, which may be used in thepresent invention. FIG. 1A is a perspective representation of anassembled sensor strip 100 including a sensor base 110, at leastpartially covered by a lid 120 that includes a vent 130, a samplecoverage area 140, and an input end opening 150. A partially-enclosedvolume 160 (the capillary gap or cap-gap) is formed between the base 110and the lid 120. Other sensor strip designs compatible with the presentinvention also may be used, such as those described in U.S. Pat. Nos.5,120,420 and 5,798,031.

A liquid sample for analysis may be transferred into the cap-gap 160 byintroducing the liquid to the opening 150. The liquid fills the cap-gap160 while expelling the previously contained air through the vent 130.The cap-gap 160 may contain a composition (not shown) that assists inretaining the liquid sample in the cap-gap. Examples of suchcompositions include water-swellable polymers, such as carboxymethylcellulose and polyethylene glycol; and porous polymer matrices, such asdextran and polyacrylamide.

FIG. 1B depicts a top-view of the sensor strip 100, with the lid 120removed. Conductors 170 and 180 may run under a dielectric layer 190from the opening 150 to a working electrode 175 and a counter electrode185, respectively. In one aspect, the working and counter electrodes175, 185 may be in substantially the same plane, as depicted in thefigure. In a related aspect, the working and counter electrodes 175, 185may be separated by greater than 200 or 250 μm and may be separated froman upper portion of the lid 120 by at least 100 μm. In another aspect,the working and counter electrodes 175, 185 may be separated by lessthan 200 μm. The dielectric layer 190 may partially cover the electrodes175, 185 and may be made from any suitable dielectric material, such asan insulating polymer.

The counter electrode 185 may support the electrochemical activity atthe working electrode 175 of the sensor strip 100. In one aspect, thepotential to support the electrochemical activity at the workingelectrode 175 may be provided to the sensor system by forming thecounter electrode 185 from an inert material, such as carbon, andincluding a soluble redox species, such as ferricyanide, within thecap-gap 160. In another aspect, the potential at the counter electrode185 may be a reference potential achieved by forming the counterelectrode 185 from a redox pair, such as Ag/AgCl, to provide a combinedreference-counter electrode. Alternatively, the sensor strip 100 may beprovided with a third conductor and electrode (not shown) to provide areference potential to the sensor system.

FIG. 2A depicts an end-view diagram of the sensor strip depicted in FIG.1B showing the layer structure of the working electrode 175 and thecounter electrode 185. The conductors 170 and 180 may lie directly onthe base 110. Surface conductor layers 270 and 280 optionally may bedeposited on the conductors 170 and 180, respectively. The surfaceconductor layers 270, 280 may be made from the same or from differentmaterials.

The material or materials used to form the conductors 170, 180 and thesurface conductor layers 270, 280 may include any electrical conductor.Preferable electrical conductors are non-ionizing, such that thematerial does not undergo a net oxidation or a net reduction duringanalysis of the sample. The conductors 170, 180 preferably include athin layer of a metal paste or metal, such as gold, silver, platinum,palladium, copper, or tungsten. The surface conductor layers 270, 280preferably include carbon, gold, platinum, palladium, or combinationsthereof. If a surface conductor layer is not present on a conductor, theconductor is preferably made from a non-ionizing material.

The surface conductor material may be deposited on the conductors 170,180 by any conventional means compatible with the operation of thesensor strip, including foil deposition, chemical vapor deposition,slurry deposition, and the like. In the case of slurry deposition, themixture may be applied as an ink to the conductors 170, 180, asdescribed in U.S. Pat. No. 5,798,031.

The reagent layers 275 and 285 may be deposited on the conductors 170and 180, respectively. The layers are formed from at least one reagentcomposition that includes reagents and optionally a binder. The binderis preferably a polymeric material that is at least partiallywater-soluble. In one aspect, the binder may form a gel or gel-likematerial when hydrated by the sample. In another aspect, the binder mayfilter red blood cells.

Suitable partially water-soluble polymeric materials for use as thebinder may include poly(ethylene oxide) (PEO), carboxy methyl cellulose(CMC), polyvinyl alcohol (PVA), hydroxyethylene cellulose (HEC),hydroxypropyl cellulose (HPC), methyl cellulose, ethyl cellulose, ethylhydroxyethyl cellulose, carboxymethyl ethyl cellulose, polyvinylpyrrolidone (PVP), polyamino acids, such as polylysine, polystyrenesulfonate, gelatin, acrylic acid, methacrylic acid, starch, maleicanhydride salts thereof, derivatives thereof, and combinations thereof.Among the above binder materials, PEO, PVA, CMC, and HEC are preferred,with CMC being more preferred at present.

In addition to the binder, the reagent layers 275 and 285 may includethe same or different reagents. When including the same reagents, thereagent layers 275 and 285 may be the same layer. In one aspect, thereagents present in the first layer 275 may be selected for use with theworking electrode 175, while the reagents present in the second layer285 may be selected for use with the counter electrode 185. For example,the reagents in the layer 285 may facilitate the free flow of electronsbetween the sample and the conductor 180. Similarly, the reagents in thelayer 275 may facilitate the reaction of the analyte.

The reagent layer 275 may include an enzyme system specific to theanalyte that may facilitate the reaction of the analyte while enhancingthe specificity of the sensor system to the analyte, especially incomplex biological samples. The enzyme system may include one or moreenzyme, cofactor, and/or other moiety that participates in the redoxreaction with the analyte. For example, an alcohol oxidase can be usedto provide a sensor strip that is sensitive to the presence of alcoholin a sample. Such a system could be useful in measuring blood alcoholconcentrations. In another example, glucose dehydrogenase or glucoseoxidase may be used to provide a sensor strip that is sensitive to thepresence of glucose in a sample. This system could be useful inmeasuring blood glucose concentrations, for example in patients known orsuspected to have diabetes.

Enzymes for use in the enzyme system include alcohol dehydrogenase,lactate dehydrogenase, β-hydroxybutyrate dehydrogenase,glucose-6-phosphate dehydrogenase, glucose dehydrogenase, formaldehydedehydrogenase, malate dehydrogenase, and 3-hydroxysteroid dehydrogenase.Preferable enzyme systems are oxygen independent, thus not substantiallyoxidized by oxygen.

One such oxygen independent enzyme family is glucose dehydrogenase(GDH). Using different co-enzymes or co-factors, GDH may be mediated ina different manner by different mediators. Depending on theirassociation with GDH, a co-factor, such as flavin adenine dinucleotide(FAD), can be tightly held by the host enzyme, such as in the case ofFAD-GDH; or a co-factor, such as Pyrroloquinolinequinone (PQQ), may becovalently linked to the host enzyme, such as with PQQ-GDH. Theco-factor in each of these enzyme systems may either be permanently heldby the host enzyme or the co-enzyme and the apo-enzyme may bere-constituted before the enzyme system is added to the reagentcomposition. The co-enzyme also may be independently added to the hostenzyme moiety in the reagent composition to assist in the catalyticfunction of the host enzyme, such as in the cases of nicotinamideadenine dinucleotide NAD/NADH+ or nicotinamide adenine dinucleotidephosphate NADP/NADPH.

The reagent layer 275 also may include a mediator to more effectivelycommunicate the results of the analyte reaction to the surface conductor270 and/or the conductor 170. Mediators may be separated into two groupsbased on their electrochemical activity. One electron transfer mediatorsare chemical moieties capable of taking on one additional electronduring the conditions of the electrochemical reaction, while twoelectron transfer mediators are chemical moieties capable of taking ontwo additional electrons during the conditions of the reaction. Asdepicted in FIG. 2B, one electron transfer mediators can transfer oneelectron from the enzyme to the working electrode, while as depicted inFIG. 2C, two electron transfer mediators can transfer two electrons.

While other mediators may be used, two electron transfer mediators arepreferred due to their ability to transfer approximately twice as manyelectrons from the enzyme system to the working electrode for the samemolar amount of mediator in relation to one electron transfer mediators.By reducing the amount of mediator required for the sensor to perform inrelation to conventional sensor strips, the sensor strips of the presentinvention may demonstrate an increase in long-term stability. Thisstability increase may be attributable to a reduction in enzymedenaturization by the mediator during storage. The stability increasealso may be attributable to a reduction in the amount of mediatoravailable to oxidize the enzyme during storage.

Examples of one electron transfer mediators include compounds, such as1,1′-dimethyl ferrocene, ferrocyanide and ferricyanide, andruthenium(III) and ruthenium(II) hexaamine. Two electron mediatorsinclude the organic quinones and hydroquinones, such as phenathrolinequinone; phenothiazine and phenoxazine derivatives;3-(phenylamino)-3H-phenoxazines; phenothiazines; and7-hydroxy-9,9-dimethyl-9H-acridin-2-one and its derivatives. Examples ofadditional two electron mediators include the electro-active organicmolecules described in U.S. Pat. Nos. 5,393,615; 5,498,542; and5,520,786, which are incorporated herein by reference, for example.

Preferred two electron transfer mediators include3-phenylimino-3H-phenothiazines (PIPT) and 3-phenylimino-3H-phenoxazines(PIPO). More preferred two electron mediators include the carboxylicacid or salt, such as ammonium salts, of phenothiazine derivatives. Atpresent, especially preferred two electron mediators include(E)-2-(3H-phenothiazine-3-ylideneamino)benzene-1,4-disulfonic acid(Structure I), (E)-5-(3H-phenothiazine-3-ylideneamino)isophthalic acid(Structure II), ammonium(E)-3-(3H-phenothiazine-3-ylideneamino)-5-carboxybenzoate (StructureIII), and combinations thereof. The structural formulas of thesemediators are presented below.

In another respect, preferred two electron mediators have a redoxpotential that is at least 100 mV lower, more preferably at least 150 mVlower, than ferricyanide.

The reagent layers 275, 285 may be deposited by any convenient means,such as printing, liquid deposition, or ink-jet deposition. In oneaspect, the layers are deposited by printing. With other factors beingequal, the angle of the printing blade may inversely affect thethickness of the reagent layers. For example, when the blade is moved atan approximately 82° angle to the base 110, the layer may have athickness of approximately 10 μm. Similarly, when a blade angle ofapproximately 62° to the base 110 is used, a thicker 30 μm layer may beproduced. Thus, lower blade angles may provide thicker reagent layers.In addition to blade angle, other factors, such as the viscosity of thematerial being applied as well as the screen-size and emulsioncombination, may affect the resulting thickness of the reagent layers275, 285.

When thinner reagent layers are preferred, deposition methods other thanprinting, such as micro-pipetting, ink jetting, or pin-deposition, maybe required. These deposition methods generally give the dry reagentlayers at micrometer or sub-micrometer thickness, such as 1-2 μm. Forexample, pin-deposition methods may provide average reagent layerthicknesses of 1 μm. The thickness of the reagent layer resulting frompin-deposition, for example, may be controlled by the amount of polymerincluded in the reagent composition, with higher polymer contentproviding thicker reagent layers. Thinner reagent layers may requireshorter pulse widths than thicker reagent layers to maintain the desiredmeasurement performance and/or substantially measure analyte within thediffusion barrier layer (DBL).

The working electrode 175 also may include a DBL that is integral to areagent layer 275 or that is a distinct layer 290, such as depicted inFIG. 2A. Thus, the DBL may be formed as a combination reagent/DBL on theconductor, as a distinct layer on the conductor, or as a distinct layeron the reagent layer. When the working electrode 175 includes thedistinct DBL 290, the reagent layer 275 may or may not reside on the DBL290. Instead of residing on the DBL 290, the reagent layer 275 mayreside on any portion of the sensor strip 100 that allows the reagent tosolubilize in the sample. For example, the reagent layer 175 may resideon the base 110 or on the lid 120.

The DBL provides a porous space having an internal volume where ameasurable species may reside. The pores of the DBL may be selected sothat the measurable species may diffuse into the DBL, while physicallylarger sample constituents, such as RBCs, are substantially excluded.Although conventional sensor strips have used various materials tofilter RBCs from the surface of the working electrode, a DBL provides aninternal porous space to contain and isolate a portion of the measurablespecies from the sample.

When the reagent layer 275 includes a water-soluble binder, any portionof the binder that does not solubilize into the sample prior to theapplication of an excitation may function as an integral DBL. Theaverage initial thickness of a combination DBL/reagent layer ispreferably less than 16 or 8 micrometers (μm) and more preferably lessthan 4 μm. At present, an especially preferred average initialthicknesses of a combination DBL/reagent layer is from 0.25 to 3 μm orfrom 0.5 to 2 μm. The desired average initial thickness of a combinationDBL/reagent layer may be selected for a specific excitation length onthe basis of when the diffusion rate of the measurable species from theDBL to a conductor surface, such as the surface of the conductor 170 orthe surface of the surface conductor 270 from FIG. 2A, becomesrelatively constant. In one aspect, the DBL/reagent layer may have anaverage initial thickness of 1 μm or less when combined with anexcitation pulse width of 0.25 seconds or less.

The distinct DBL 290 may include any material that provides the desiredpore space, while being partially or slowly soluble in the sample. Inone aspect, the distinct DBL 290 may include a reagent binder materiallacking reagents. The distinct DBL 290 may have an average initialthickness of from 1 to 15 μm, and more preferably from 2 to 5 μm.

FIG. 3 represents an electrochemical analysis 300 for determining thepresence and optionally the concentration of an analyte 322 in a sample312. In 310, the sample 312 is introduced to a sensor strip 314, such asthe sensor strip depicted in FIGS. 1A-1B and 2A. The reagent layers,such as 275 and/or 285 from FIG. 2A, begin to solubilize into the sample312, thus allowing reaction. At this point in the analysis, it may bebeneficial to provide an initial time delay, or “incubation period,” forthe reagents to react with the sample 312. Preferably, the initial timedelay may be from 0.5 to 5 seconds. A more in-depth treatment of initialtime delays may be found in U.S. Pat. Nos. 5,620,579 and 5,653,863.

During the reaction, a portion of the analyte 322 present in the sample312 is chemically or biochemically oxidized or reduced in 320, such asby an oxidoreductase. Upon oxidation or reduction, electrons may betransferred between the analyte 322 and a mediator 332 in 330, such asthrough the oxidoreductase.

In 340, the charged mediator 332 is electrochemically excited (oxidizedor reduced). For example, when the sample 312 is whole blood containingglucose oxidized by the PQQ-GDH enzyme system in 320, which thentransfers two electrons to reduce a phenothiazine derivative mediator in330, the excitation of 340 oxidizes the phenothiazine derivativemediator at the working electrode. In this manner, electrons areselectively transferred from the glucose analyte to the workingelectrode of the sensor strip where they may be detected by ameasurement device.

The current resulting from the excitation 340 may be recorded during theexcitation 340 as a function of time in 350. In 360, the sampleundergoes relaxation. Preferably, the current is not recorded during therelaxation 360. The recorded current and time values may be analyzed todetermine the presence and/or concentration of the analyte 322 in thesample 312 in 370.

Amperometric sensor systems apply a potential (voltage) to the sensorstrip to excite the measurable species while the current (amperage) ismonitored. Conventional amperometric sensor systems may maintain thepotential while measuring the current for a continuous read pulse lengthof from 5 to 10 seconds, for example. In contrast to conventionalmethods, the duty cycles used in the electrochemical analysis 300replace continuous, long-duration read pulses with multiple excitationsand relaxations of short duration. A more detailed description ofmultiple excitation and relaxation or “gated” pulse sequences may befound in PCT/US2006/028013, filed Jul. 19, 2006, entitled “GatedAmperometry.”

Referring to FIG. 3, the excitation 340, the recordation 350, and therelaxation 360 constitute a single duty cycle. Preferably, at least 2,4, 6, or 7 duty cycles are applied during an independently selected 3,5, 7, 9, or 14 second time period. In one aspect, the duty cycles areapplied during a 3 to 14 second time period. In another aspect, at least4 duty cycles are applied within 30 seconds, 9 seconds, or less. Inanother aspect, from 2 to 6 duty cycles may be applied within 10 secondsor less. In another aspect, from 2 to 4 duty cycles may be appliedwithin 3 to 9 seconds.

After the excitation 340, in 360 the measurement device may open thecircuit through the sensor strip 314, thus allowing the system to relax.During the relaxation 360, the current present during the excitation 340is substantially reduced by at least one-half, preferably by an order ofmagnitude, and more preferably to zero. Preferably, a zero current stateis provided by an open circuit or other method known to those ofordinary skill in the art to provide a substantially zero current flow.In one aspect, the relaxation 360 is at least 0.5 or at least 0.2seconds in duration. In another aspect, the relaxation 360 is from 0.2to 3 seconds or from 0.5 to 1 second in duration.

FIGS. 4A-4D depict examples of gated amperometric pulse sequences wheremultiple duty cycles were applied to the sensor strip after introductionof the sample. In these examples, square-wave pulses were used; however,other wave types compatible with the sensor system and the test samplealso may be used.

Each depicted pulse sequence includes an initial one second excitationpulse 420 followed by multiple 0.25 second excitations 430. The initialexcitation pulse 420 may be of longer duration than the subsequentexcitations 430. For example, the initial excitation pulse 420 may rangefrom 0.75 to 3 seconds in duration. The length of the excitation pulse420 may be tailored to oxidize the relatively small amount of enzymesystem deposited on the strip. When the initial excitation pulse 420 isused, it is preferably of longer duration than the following multipleexcitations 430. For example, the pulse sequence may include the initialexcitation pulse 420 having a duration of 2 seconds, followed by an opencircuit relaxation of 3 seconds, followed by the subsequent excitation430 having a duration of 0.125 seconds.

The multiple excitations 430 may range from 0.01 to 3 seconds induration. In one aspect, the total excitation length is two seconds orless, thus including the initial one second pulse 420 followed by fourof the 0.25 second excitations 430. In another aspect, the totalexcitation length is 1.5 seconds or less, thus including the initial onesecond pulse 420 followed by one or two of the 0.25 second excitations430. The pulse sequence may include additional excitations, such asexcitation 440 depicted in FIG. 4A. In another aspect, the excitationsmay be of different amplitudes, such as depicted in FIGS. 4C and 4D. Ina preferred aspect, when excitations of different amplitudes are used,the difference in amplitudes may be within 500 mV.

The short excitations may permit the accurate analysis of the samplewith reagent compositions having reduced polymer, enzyme system, andmediator concentrations in relation to conventional compositions.Furthermore, the short excitations allow for the analysis to becompleted within 8.5 seconds or less, or more preferably, 5 seconds orless from the initial application of a signal to the sensor strip.

Either short or long excitations may be used. Preferably, the currentmeasurement that the analyte concentration is determined from is takenwithin 2 seconds or 1 second of the initial application of the signal.More preferably, multiple short excitations are combined with a currentmeasurement taken within 2 seconds, 1 second, or less from the initialapplication of the signal to determine the analyte concentration of thesample.

In combination with the gated amperometric pulse sequences of thepresent invention, reagent compositions including specific amounts ofpolymeric binder, enzyme system, and/or mediator were found to reducehematocrit bias and/or increase long-term stability. While conventionalstrips are often described in terms of the percent of each reagentcomposition ingredient, it is the density of each reagent compositioningredient (absolute amount per area) that is relevant to long-termstability. By limiting the amount of mediator and enzyme available forinteraction, the amount of environmentally reduced mediator (mediatornot responsive to the underlying analyte concentration) present at thetime of analysis may be substantially reduced. Thus, providing abeneficial reduction in background current. As conventional biosensorsuse excess enzyme in an attempt to improve long-term stability, it wasunexpected that a reduction in the enzyme system and/or mediatorprovided an improvement in long-term stability for the presentbiosensors. Furthermore, if individual reagent compositions areoptimized for each electrode, the relative amount of mediator availableto interact with the enzyme may be further reduced in relation to singlereagent composition sensor strips.

For example, hematocrit biases were larger with the higher polymerconcentrations of 2%, 1.5%, and 1% than with a 0.5% (w/w) polymerconcentration in the reagent solution when using high enzyme loading.This may be attributable to higher polymer concentrations producingthicker reagent layers that affect hematocrit bias on re-hydration.Thus, as assay times become shorter, rapid re-hydration without anincrease in hematocrit bias may be preferred. In this manner, apreferable balance may be reached between the polymer content and enzymeloading of the reagent composition to provide the desired level ofhydration during the time of the assay. When applied to the sensorstrip, polymer deposition densities of 2 μg/mm² or less may bepreferred, with polymer deposition densities from 0.8 to 1.5 μg/mm²being more preferred.

When different deposition solutions are deposited on the working andcounter electrodes, the amount of the enzyme system present on theworking electrode is controlling. Thus, depositing approximately 0.24 μLof a solution including 1.2 U of the PQQ-GDH enzyme system, 2.57 μg ofthe Structure I mediator, 1.18 μg of the CMC polymer, and 3.28 μg ofphosphate buffer provides a reagent solution concentration of 5 U/μLenzyme system, 24 mM mediator, 0.5% polymer, and 100 mM phosphatebuffer.

Depositing this reagent solution on a deposition area of 1.5 mm²including a working electrode area of 1 mm², provides an enzyme systemdeposition density at the working electrode of about 0.8 U/mm² (1.2U/1.5 mm²), a mediator deposition density of 1.72 μg/mm² (2.57 μg/1.5mm²), a polymer deposition density of 0.8 μg/mm², and a phosphate bufferdeposition density of 2.2 μg/mm². Similarly, for a reagent solutionincluding approximately 2 U/μL of the PQQ-GDH enzyme system, 24 mM ofthe Structure I mediator, 0.5% CMC, and 100 mM phosphate buffer, theenzyme system deposition density will be about 0.3 U/mm². Furthermore,the specific activity of an enzyme system may be translated into weightdensity in terms of μg/mm². For example, if the activity of an enzymesystem is 770 U/mg, the enzyme system's activity density of 0.3 U/mm²becomes 0.39 μg/mm² in weight density. Thus, the activity of an enzymesystem may be taken into account when determining the deposition densityfor the enzyme system.

Conventional reagent layers include enzyme deposition densities of 1 to6 μg/mm². In contrast, after deposition, reagent compositions of thepresent invention may include enzyme system deposition densities of lessthan 1 μg/mm², preferably less than 0.5 μg/mm². The deposition solutionsof the present invention may include about 4 U/μL or less of the enzymesystem or may include about 3 U/μL or less of the enzyme system. Atpresent, 2.2 U/μL or less of the PQQ-GDH enzyme system may be includedin the reagent solution. Thus, when the reagent composition is appliedto the sensor strip, 1.3 Units or less of the enzyme system may bepresent on the sensor strip, with from 0.3 to 0.8 Units being morepreferred.

Conventional sensor strips generally have mediator deposition densitiesranging from 10 to 25 μg/mm². In contrast, mediator deposition densitiesof 8 μg/mm² and less are preferred for the present invention, withmediator deposition densities of 5 μg/mm² and less being more preferred.At present, mediator deposition densities of 2 μg/mm² and less areespecially preferred. In a preferred aspect, two electron mediators arepreferable to one electron mediators.

Table I, below, provides the deposition densities of the polymericbinder, buffer, enzyme system, and mediator components in the reagentcompositions used in FIGS. 5A-5E and FIGS. 6A-6C, discussed below. Theenzyme system deposition density for reagent composition 3 (RC3) at 0.42μg/mm² is nearly 60% less than the lowest 1 μg/mm² value of thepreviously discussed conventional strips.

Similarly, the mediator densities given in Table I are approximately oneorder of magnitude smaller than those of conventional sensor strips. Forexample, the mediator deposition density of RC3 was approximatelyone-fifth that of a conventional sensor strip. This reduction in thedeposition density of the mediator in relation to conventional sensorstrips may provide a substantial increase in the long-term stability ofthe reagent compositions of the present invention.

TABLE I CMC, Buffer, PQQ/GDH, Mediator, μg/mm² μg/mm² μg/mm² μg/mm² RC11.59 2.17 3.99 1.72 RC2 0.79 2.19 1.07 1.72 RC3 0.79 2.19 0.42 1.72 RC40.77 2.22 0.99 1.36

FIG. 5A depicts the glucose dose response curves at different hematocritlevels in WB samples for the RC2 sensor strip that included 1.2 unit persensor (U/sensor) at a deposition density of 0.8 U/mm² of the PQQ-GDHenzyme. FIG. 5B depicts the glucose dose response curves at differenthematocrit levels in WB samples for the RC3 sensor strip that included0.5 U/sensor at a deposition density of 0.3 U/mm² of the PQQ-GDH enzyme.Both the 1.2 U and 0.5 U sensors delivered hematocrit biases less than±5%. However, a difference between the 1.2 U and 0.5 U sensors wasobserved with regard to the system sensitivity and intercept.

A reduction in the ratio of intercept-to-slope may be seen for the 0.5U/sensor of FIG. 5B by comparing the intercept to slope (I/S) ratios,expressed in units of mg/dL, with those of the 1.2 U/sensor of FIG. 5A,with lower ratios representing a reduction in background signal. Thus,the I/S ratio for the 1.2 U/sensor of FIG. 5A is ˜6 mg/dL, while the I/Sratio for the 0.5 U/sensor of FIG. 5B is ˜0.15 mg/dL, an approximate 40time reduction at the 40% hematocrit concentration. The superiorbackground signal performance of the 0.5 U sensor in relation to a 1.5 Usensor was established.

In one aspect, the performance characteristic of low background (lowintercept) may be provided by relatively low mediator and enzymedeposition densities. In another aspect, the performance characteristicof high sensitivity (high slope) may be provided by optimizing thetiming ratio of the excitations and relaxations of the pulse sequence.For example, combining a relatively long relaxation period of 1.5seconds with a relatively short excitation period of 0.25 secondprovides a large current density at the surface of the workingelectrode. Numerically, the relatively small intercept value over therelatively large slope value further improves the I/S. Thus, preferablesensor performance characteristics (small I/S values) may be provided bythe reagent composition and the measurement method in combination.

Equation (1), below, provides an analytical relationship between thecurrent imprecision and the resulting imprecision in the determinedglucose concentration as a function of the I/S ratio when the glucoseand current relationship is i=S*G+Int:

$\begin{matrix}{{\frac{\Delta\; G\text{/}G}{\Delta\; i\text{/}i} = {\frac{{SD}_{G}\text{/}G}{{SD}_{i}\text{/}i} = {\frac{\%\mspace{14mu}{CV}_{G}}{\%\mspace{14mu}{CV}_{i}} = {1 + {\frac{1}{G}\frac{Int}{S}}}}}},} & (1)\end{matrix}$where G denotes the glucose concentration, ΔG/G is the relative error ofthe glucose concentration in mg/dL, Δi/i is the relative error of themeasured current, SD_(G)/G is the relative standard deviation in ΔG/G,SD_(i)/i is the relative standard deviation in Δi/i, % CV_(G) is thecoefficient of variance, which is proportional to relative standarddeviation and represents the glucose measurement precision, % CV_(i) isthe coefficient of variance, which is proportional to relative standarddeviation and represents the current measurement precision, and Int/S isthe intercept-to-slope (I/S) ratio in mg/dL. Equation (1) may be derivedfrom i=S*G+Int by taking the derivative of the reverse functionG=f(i)=(i−Int)/S in the derivation and analysis of error propagation.Due to the 1+1/G term, a current imprecision of 1 results in a glucoseconcentration imprecision of greater than 1 if the I/S value is greaterthan zero. Intercept to slope (I/S) ratios expressed in mg/dL may beexpressed in terms of millimoles/Liter (mM/L) by dividing the I/S(mg/dL) value by 18 ([mg/dL]/18=[mM] for glucose).

Current imprecision describes the variance between the currentmeasurements of multiple sensor strips. Thus, current imprecisionrepresents the amount that a recorded current value differs from themean current value when identical glucose samples are analyzed usingmultiple sensor strips. The more a recorded current value from aparticular strip deviates from the mean value for multiple strips, thepoorer the current recorded from that strip correlates with the actualglucose concentration of the sample. Thus, the current measurementprecision may measured by its imprecision or %-CV_(i), and the glucosemeasurement precision also may be measured by its imprecision or%-CV_(G).

Table II, below, provides slopes (1/G values) calculated with Equation(1) from multiple glucose concentrations in mg/dL. Thus, 1/57.6=0.0174,1/111=0.009, 1/222.25=0.0045, 1/444.75=0.0022, and 1/669=0.0015, wherethe denominators are the plasma glucose concentrations tested. As theglucose concentration increases, the slope value decreases according toequation (1).

TABLE II Glucose I/S Ratio Experimentally Concentration Calculated fromDetermined in mg/dL Equation (1) I/S Ratio 57.6 0.0174 0.0174 111 0.0090.009 222.25 0.0045 0.0045 444.75 0.0022 0.0022 669 0.0015 0.0015

FIG. 5C presents I/S ratios in mg/dL determined from experimental data.The current values were obtained with gated amperometric pulse sequencesfrom WB samples including 40% hematocrit and glucose concentrations of57.6, 111, 222.25, 444.75, or 669 mg/dL. Calibration constants weredetermined at 4, 5.5, 7, 8.5, 10, 11.5, 13, and 14.5 seconds into theanalysis. As can be seen from the table, the calculated andexperimentally determined values are the same, confirming the ability ofEquation (1) to describe the behavior of the sensor strip regardingglucose and current imprecision.

For low glucose concentrations, such as 50 mg/dL and below, it ispreferable to maintain a low I/S value to reduce the imprecisionintroduced into the determined glucose concentration from the underlyingimprecision in the recorded current value. For example, if the I/S ratiois 50 mg/dL and the glucose concentration is 50 mg/dL, then the ratio of%-CV_(G)/%-CV_(i) is 2 [1+(I/S)/G=1+50/50=2], and any currentimprecision will be amplified by a factor of 2 in the determined glucoseconcentration. Thus, if the typical imprecision in the recorded currentvalues is 3.5%, a system having an I/S ratio of 50 mg/dL results in aglucose imprecision of 7%.

FIG. 5D shows I/S ratios from 0 to 20 mg/dL for glucose concentrationsof 55.375, 112.25, and 433.5 mg/dL determined from sensor strips havingreagent compositions RC2, RC3, or RC4 from Table I, above. As the I/Sratios decrease from 20, the imprecision associated with the determinedglucose concentration also decrease. Thus, demonstrating that lower I/Sratios are preferred at lower glucose concentrations, as previouslydiscussed.

When each of these factors are considered, sensor strips having an I/Sratio of 20 mg/dL or less are preferred, with those having I/S ratios of10 mg/dL or less or 6 mg/dL or less being more preferred. At present,strips having an I/S ratio of 1 mg/dL or less are even more preferred.

FIG. 5E shows the nearly identical hematocrit performance of a sensorsystem in accord with the present invention used with plasma and 40%hematocrit whole blood samples. In this example, the reagent compositionof the sensor strip was RC2, as described in Table I, above. Thus, thegated pulse sequences in combination with the reagent compositions ofthe present invention provide a substantial reduction in the hematocriteffect observed for WB samples.

FIG. 6A shows the effect of environmental stress on the long-termstability of sensor strips. The Y-axis of the graph shows the absolutestability bias in terms of mg/dL for glucose concentrations below 75mg/dL or the %-bias for glucose concentrations at and above 75 mg/dL forenvironmentally stressed sensor strips in relation to sensor stripsstored at −20° C. The stressed sensor strips had been stored at 50° C.for two weeks as an accelerated process simulating storage at 25° C. for18 months. Since the mediator density was initially low (1.72 μg/mm²),the average stability bias line in FIG. 6A established that as theenzyme density decreased from RC1 (3.2 U/mm², or 4 ug/mm² at 770 U/mgspecific enzyme activity) to RC2, RC3 and RC4 (both at 0.8 U/mm² orless), so did the environmentally induced biases.

FIG. 6A also established that a reagent composition in accord with thepresent invention may provide long-term stability sufficient to notrequire an initial long oxidative pulse before the analysis, such asthat described in U.S. Pat. No. 5,653,863 to Genshaw et al. The analysisthat generated the data of FIG. 6A had a total duration of 4.5 secondswith a total oxidation time of 1.5 seconds (corresponding to pulsesequence of FIG. 4A). Thus, only 30% of the assay time was spentoxidizing the WB sample in FIG. 6A, compared to ˜67% of the total assaytime in the U.S. Pat. No. 5,653,863 analysis (20 seconds of oxidationover a 30 second total analysis time). In one aspect, 45% or less of thetotal assay time is spent oxidizing the sample. In another aspect, 35%or less of the total assay time is spent oxidizing the sample.

While the Structure III mediator appeared to have a slightly lowerstability bias than the Structure I mediator in FIG. 6A, this smalldifference may be attributable to other factors within the test system.Thus, reagent compositions RC2, RC3, and RC4 in combination with thegated pulse sequences of the present invention demonstrated a stabilitybias of less than ±5% after exposure to 50° C. temperatures for twoweeks, while RC1 with a higher enzyme system deposition density did not.

FIG. 6B shows the effect of environmental stress on the long-termstability of sensor strips stored at 50° C. for four weeks. The Y-axisof the graph shows the absolute stability bias in terms of mg/dL forglucose concentrations below 75 mg/dL or the %-bias for glucoseconcentrations at and above 75 mg/dL for environmentally stressed sensorstrips in relation to sensor strips stored at −20° C. The increase instability bias for the 4 week storage period of FIG. 6B in relation tothe shorter 2 week storage period of FIG. 6A is seen in the Y-axisvalues for RC1. Since the mediator density was initially low (1.72ug/mm²), the average stability bias line in FIG. 6B established that asthe enzyme density decreased from RC1 (3.2 U/mm², or 4 ug/mm² at 770U/mg specific enzyme activity) to RC2, RC3, and RC4 (both at 0.8 U/mm²or less), so did the environmentally induced biases. Thus, the averagestability bias line for the longer 4 week storage period of FIG. 6B alsodecreases substantially for RC2, RC3, and RC4 in relation to RC1. Evenafter 4 weeks of storage at 50° C., reagent compositions RC2, RC3, andRC4 in combination with the gated pulse sequences of the presentinvention demonstrated a stability bias of less than ±5%.

FIG. 6C shows the stability bias for three reagent compositionsincluding the PQQ-GDH enzyme system after 52 weeks at 25° C. under 60%relative humidity. Unlike in FIGS. 6A and 6B a longer time period at alower temperature was used to age the strips. The bias values for RC2and RC3 were similar to those observed under accelerated aging, beingless than ±5% on average. RC4 showed an increase in bias in relation tothe accelerated aging results of FIGS. 6A and 6B, increasing to a bitless than the +10% level on average. This increase may be attributableto experimental error or to a stability issue with the lot of StructureIII mediator used in RC4 under the specific circumstances of this test.

FIG. 6D shows the stability bias values for sensor strips including RC2where the data point used to calculate the analyte concentration wastaken at a specific time after starting the analysis. For example, the2.75 second line represents the bias/%-bias values of analyteconcentration as determined after the first short excitation pulse, suchas the excitation 430 in FIG. 4A. Similarly, the 4.375 second linerepresents the bias/%-bias values of analyte concentration as determinedafter a second short excitation pulse. Thus, using an environmentallystressed sensor strip including RC2 and a gated pulse sequence, theanalyte concentration of the sample may be accurately determined in lessthan 3 seconds.

Example 1: Sensor Strip Preparation

In one aspect, sensor strips in accord with the present invention weremade using pin-deposition. The counter electrode reagent solution wasprepared by making a stock solution of 100 mM phosphate buffer in 1%carboxylmethyl cellulose (CMC). Then, enough powdered mediator wasdissolved into the buffer/CMC solution to make a final solution of 100mM mediator in 100 mM phosphate buffer and 1% CMC.

The working electrode reagent solution was prepared typically by makinga stock solution of 100 mM phosphate buffer in 0.5% CMC (100 mL forexample). Then, enough powdered mediator was dissolved into thebuffer/CMC stock solution to make a solution (10 mL for example) of24-25 mM mediator in 100 mM phosphate buffer and 0.5% CMC. Finally, foran enzyme loading of 5 U/μL, about 33.4 mg of the PQQ-GDH enzyme system(749 U/mg specific activity) was combined with 5 mL of themediator/buffer/CMC solution in a glass container ([5000 U/mL*5 mL]/[749U/mg]=33.4 mg). The mixture was slowly swirled to dissolve the dryenzyme powder into the solution. This formulation was for RC2. If only 2mL of final reagent solution was needed, then about 13.4 mg PQQ-GDHenzyme with 749 U/mg specific activity was weighed out. Similarly forRC3 composition of 2 U/μL, about 13.4 mg of PQQ-GDH enzyme with thespecific activity of 749 U/mg was dissolved by 5 mL of themediator/buffer/CMC solution to make up the final reagent solution fordeposition. If only 2 mL of the final reagent solution was needed, thenabout 5.34 mg PQQ-GDH enzyme at 749 U/mg specific activity was weighedout to make the final solution.

Pin-deposition was used to deposit the counter electrode reagentsolution on one of the carbon surfaces to form the counter electrode.The volume delivered by each deposition was from about 0.2 to 0.24 μL,which spread to cover an area of about 1.5 to 2 mm². Similarly,pin-deposition was used to deposit the working electrode reagentsolution on one of the carbon surfaces to form the working electrode.The volume for each working electrode deposition was also about 0.2-0.24μL with a similar solution spread over the carbon to form the workingelectrode. The completed sensor sheet was allowed to air-dry for 15minutes followed by storage in a desiccated container before finallamination to form the completed sensor strips. In one aspect, reagentsolution deposition densities of 0.16 μL/mm² (0.24 uL/1.5 mm²) or lessat the working electrode are preferred.

Example 2: Donor Study

Whole blood samples were collected from 21 subjects having diabetesmellitus. Two analyses were performed on each subject to provide 42measurements from each of the four sensor strip lots SS1 through SS4 asprovided in Table III, below. Different working and counter electrodereagent compositions were used for SS1 through SS3, while SS4 used asingle reagent composition substantially covering both conductors. SS1,SS2, and SS3 represent different manufacturing lots using the samereagent amounts. The volume of reagent composition deposited on theindividual conductors for SS1 through SS3 was about 0.24 μL, with twodeposition areas of about 1.5 mm² each. The volume of reagentcomposition deposited to cover both conductors for SS1 was about 0.3 μL,with a total deposition area of about 3 mm².

TABLE III Counter Electrode Working Electrode Deposition Deposition SS11.62 μg/mm² CMC 1.62 μg/mm² CMC 1.11 μg/mm² Buffer 1.11 μg/mm² Buffer0.58 μg/mm² PQQ/GDH Enzyme 1.39 μg/mm² Structure I System Mediator 1.39μg/mm² Structure I Mediator pH 7.2 ± 0.1 pH 7.2 ± 0.1 SS2 1.62 μg/mm²CMC 1.62 μg/mm² CMC 1.11 μg/mm² Buffer 1.11 μg/mm² Buffer 0.58 μg/mm²PQQ/GDH Enzyme 1.39 μg/mm² Structure I System Mediator 1.39 μg/mm²Structure I Mediator pH 7.2 ± 0.1 pH 7.2 ± 0.1 SS3 1.62 μg/mm² CMC 1.62μg/mm² CMC 1.11 μg/mm² Buffer 1.11 μg/mm² Buffer 0.58 μg/mm² PQQ/GDHEnzyme 1.39 μg/mm² Structure I System Mediator 1.39 μg/mm² Structure IMediator pH 7.2 ± 0.1 pH 7.2 ± 0.1 SS4 1.14 μg/mm² CMC Same as WorkingElectrode 0.78 μg/mm² Buffer 0.41 μg/mm² PQQ/GDH Enzyme System 1.95μg/mm² Structure I Mediator pH 7.2 ± 0.1

The samples were analyzed with a pulse sequence having an initialexcitation including two short excitations separated by about 0.25second followed by a 1 second relaxation. After the initial excitationand relaxation, a sequence of three about 0.375 second excitationsseparated by two about 1 second relaxations was applied. An outputcurrent recorded after about 5 seconds from the application of theinitial input signal was used to determine the glucose concentration ofthe sample.

Tables IV and V, below, provide the statistical hematocrit bias resultsfor SS1 through SS4, where two analyses were performed with each type ofsensor strip for each of the 21 blood samples to provide 42 readings.Table IV shows the slope, intercept and I/S ratio for each sensor striplot, while Table V shows the percentage of readings having biases within±15/±15%, 10/±+10%, or ±5/±5% of the YSI reference value. For glucoseconcentrations less than 75 mg/dL bias is expressed as mg/dL (absolute)and for glucose concentrations of 75 mg/dL and greater bias is expressedin percent (relative).

TABLE IV Sensor Slope Intercept I/S Ratio mg/dL SS1 34.61 91.73 2.65 SS232.65 −115.8 −3.55 SS3 30.76 −28.00 −0.91 SS4 33.91 106.87 3.15

TABLE V Bias % within Bias % within Sensor ±15% ±10% Bias % within ±5%SS1 100 100 93 SS2 100 97.6 64 SS3 100 97.6 93 SS4 100 97.6 67

Table IV establishes that each of the sensor strip lots had an I/S ratioof less than 5 mg/dL, thus establishing the superior background signalperformance of the strips. The performance of a sensor system may becharacterized by the spread of the biases against the reference values.This spread may be measured by the percentage of bias values fallingwithin certain limits, such as ±15 mg/dL/±15% or ±10 mg/dL/±10%. Thesmaller the limit, the better the performance. Multiple factors,including measurement imprecision and the hematocrit effect, willcontribute to the bias values. Normally, the performance is judged byhaving at least 95% of the data population being within a certainperformance limit. Thus, Table V establishes that 100% of the datapopulation from SS1 through SS4 was within the limit of ±15 mg/dL/±15%.Furthermore, more than 95% of the data population from SS1 through SS4was within the limit of ±10 mg/dL/±10%.

FIG. 7A depicts the dose response curve from the SS1 lot for the 21whole blood samples. The R² value of 0.997 established the ability ofthe strip to provide current values accurately reflecting the actualglucose concentration of the samples. FIG. 7B depicts the bias spreadsfor lots SS1 through SS4 with the 42 readings from 21 samples across theglucose concentration range in the donor study. The figure establishesthat 100% of the bias values for all of the sensor strips fell within±15%, and that the performance of the SS1 and SS3 manufacturing lots wassuperior under the conditions of the study with 93% of the datapopulation being within the ±5% limit. The narrow bias spread for thesensor strips may be attributed to the enhanced sensitivity andprecision, reflected in the small I/S ratios, and to the smallhematocrit effect provided by the strips.

FIG. 7C depicts the hematocrit sensitivity, hematocrit bias in relationto hematocrit content of the sample, obtained from lot SS1 in comparisonto a conventional strip having a deposition density of 2.96 μg/mm² for aHEC polymer, 0.69 μg/mm² for a citric buffer, 2.14 μg/mm² for a glucoseoxidase enzyme, and 13 μg/mm² for the ferricyanide mediator. Ahematocrit sensitivity of −0.26 was obtained for the SS1 lot, while theconventional sensor strips had a slope of −1.26, with numerically largerslope values indicating greater hematocrit sensitivity. Thus, thehematocrit sensitivity of the SS1 lot was approximately 79% less thanthat provided by the conventional strips.

To provide a clear and consistent understanding of the specification andclaims, the following definitions are provided.

“System” is defined as an electrochemical sensor strip in electricalcommunication through its conductors with an electronic measurementdevice, which allows for the quantification of an analyte in a sample.

“Measurement device” is defined as an electronic device that can applyan electrical input signal and measure the resulting output signal. Themeasurement device also may include the processing capability todetermine the presence and/or concentration of one or more analytes inresponse to the output signal.

“Analyte” is defined as one or more substances present in a sample. Ananalysis determines the presence and/or concentration of the analytepresent in the sample.

“Sample” is defined as a composition that may contain an unknown amountof the analyte. Typically, a sample for electrochemical analysis is inliquid form, and preferably the sample is an aqueous mixture. A samplemay be a biological sample, such as blood, urine, or saliva. A samplealso may be a derivative of a biological sample, such as an extract, adilution, a filtrate, or a reconstituted precipitate.

“Conductor” is defined as an electrically conductive substance thatremains stationary during an electrochemical analysis. Examples ofconductor materials include solid metals, metal pastes, conductivecarbon, conductive carbon pastes, and conductive polymers.

“Non-ionizing material” is defined as a material that does not ionizeduring the electrochemical analysis of an analyte. Examples ofnon-ionizing materials include carbon, gold, platinum and palladium.

“Measurement performance” is defined in terms of accuracy and/orprecision. Thus, an increase in measurement performance may be anincrease in accuracy and/or precision of the measurement.

“Precision” is defined as how close multiple analyte measurements arefor the same sample. Precision may be expressed in terms of the spreador variance among multiple measurements in relation to a mean.

“Accuracy” is defined as how close the amount of analyte measured by asensor strip corresponds to the true amount of analyte in the sample.Accuracy may be expressed in terms of bias, with larger bias valuesreflecting less accuracy.

“Bias” is defined as the difference between a measured value and theaccepted reference value. Bias may be expressed in terms of “absolutebias” or “relative bias”. Absolute bias may be expressed in the units ofthe measurement, such as mg/dL, while relative bias may be expressed asa percentage of the absolute bias value over the reference value. Eitherhematocrit or stability bias may be expressed in terms of an absolutebias value or as a percentage. Hematocrit bias uses an analyteconcentration obtained with a reference instrument, such as the YSI 2300STAT PLUS™ available from YSI Inc., Yellow Springs, Ohio, as theaccepted reference value. Stability bias uses an analyte concentrationobtained from a sensor strip stored at a temperature of −20° C. tosubstantially reduce thermal alteration of the reagent composition.

“Hematocrit sensitivity” is defined as the degree to which changes inthe hematocrit level of a sample affect the hematocrit bias values foran analysis.

“Long-term stability” is defined in relation to sensor strips packaged,such as with foil and desiccant, and stored at −20° C. for 2 or 4 weeksafter manufacture versus sensor strips exposed to 50° C. for 2 or 4weeks, respectively. Storage at 50° C. for 2 weeks may be considered toapproximate 18 months of room temperature storage. The average change ordeviation for the 0%, 50%, 100%, and 400% hematocrit levels inmeasurement performance for the 50° C. versus the −20° C. stored stripindicates the long-term stability drift or “stability bias” for thesensor strip. In this instance, an increase in background signal or biasshows a decrease in measurement performance for the sensor strip. Thus,by storing the packaged sensor strips at 50° C. and observing the biaschange in relation to −20° C. stored strips, an indication of how muchbias will increase for sensor strips remaining on store shelves forvarious time periods may be obtained.

“Mediator” is defined as a substance that may be oxidized or reduced andthat may transfer one or more electrons. A mediator is a reagent in anelectrochemical analysis and is not the analyte of interest, butprovides for the indirect measurement of the analyte. In a simplesystem, the mediator undergoes a redox reaction in response to theoxidation or reduction of the analyte. The oxidized or reduced mediatorthen undergoes the opposite reaction at the working electrode of thesensor strip and is regenerated to its original oxidation number.

“Measurable species” is defined as any electrochemically active speciesthat may be oxidized or reduced under an appropriate potential at theelectrode surface of an electrochemical sensor strip. Examples ofmeasurable species include an analyte, a substrate, or a mediator.

“Oxidoreductase” is defined as any enzyme that facilitates the oxidationor reduction of a measurable species. An oxidoreductase is a reagent.The term oxidoreductase includes “oxidases,” which facilitate oxidationreactions where molecular oxygen is the electron acceptor; “reductases,”which facilitate reduction reactions where the analyte is reduced andmolecular oxygen is not the analyte; and “dehydrogenases,” whichfacilitate oxidation reactions in which molecular oxygen is not theelectron acceptor. See, for example, Oxford Dictionary of Biochemistryand Molecular Biology, Revised Edition, A. D. Smith, Ed., New York:Oxford University Press (1997) pp. 161, 476, 477, and 560.

“Electro-active organic molecule” is defined as an organic moleculelacking a metal that is capable of undergoing a redox reaction.Electro-active organic molecules can behave as redox species and/or asmediators. Examples of electro-active organic molecules include coenzymepyrroloquinoline quinone (PQQ), benzoquinones and naphthoquinones,N-oxides, nitroso compounds, hydroxylamines, oxines, flavins,phenazines, phenothiazines, indophenols, and indamines.

“Binder” is defined as a material that provides physical support andcontainment to the reagents while having chemical compatibility with thereagents.

“Average initial thickness” is defined as the average height of a layerin its dry state prior to introduction of a liquid sample. The termaverage is used because the top surface of the layer is uneven, havingpeaks and valleys.

“Deposition density” is defined as the mass of a material deposited onan area. For example, when 0.24 μL of a solution containing 2.57 μg of asolid is deposited on a surface having an area of 1.5 mm², a depositiondensity of 1.72 μg/mm² (2.57 μg/1.5 mm²) results.

“Enzyme unit” (U) is defined as the amount of an enzyme system that willcatalyze the transformation (oxidation or reduction) of 1 micromole ofsubstrate (analyte) in 1 minute under standard conditions.

“Enzyme activity” or “activity” with regard to an enzyme system is thenumber of enzyme units per volume. Thus, activity may be given in termsof U/L or mU/mL where 1 U/L=μmol/minute/Liter, for example.

“Redox reaction” is defined as a chemical reaction between two speciesinvolving the transfer of at least one electron from a first species toa second species. Thus, a redox reaction includes an oxidation and areduction. The oxidation portion of the reaction involves the loss of atleast one electron by the first species, and the reduction portioninvolves the addition of at least one electron to the second species.The ionic charge of a species that is oxidized is made more positive byan amount equal to the number of electrons transferred. Likewise, theionic charge of a species that is reduced is made less positive by anamount equal to the number of electrons transferred.

“Oxidation number” is defined as the formal ionic charge of a chemicalspecies, such as an atom. A higher oxidation number, such as (Ill), ismore positive, and a lower oxidation number, such as (II), is lesspositive. A neutral species has an ionic charge of zero (0). Theoxidation of a species results in an increase in the oxidation number ofthat species, and reduction of a species results in a decrease in theoxidation number of that species.

“Redox pair” is defined as two conjugate species of a chemical substancehaving different oxidation numbers. Reduction of the species having thehigher oxidation number produces the species having the lower oxidationnumber. Alternatively, oxidation of the species having the loweroxidation number produces the species having the higher oxidationnumber.

“Oxidizable species” is defined as the species of a redox pair havingthe lower oxidation number, and which is thus capable of being oxidizedinto the species having the higher oxidation number. Likewise, the term“reducible species” is defined as the species of a redox pair having thehigher oxidation number, and which is thus capable of being reduced intothe species having the lower oxidation number.

“Soluble redox species” is defined as a substance that is capable ofundergoing oxidation or reduction and that is soluble in water (pH 7,25° C.) at a level of at least 1.0 grams per Liter. Soluble redoxspecies include electro-active organic molecules, organotransition metalcomplexes, and transition metal coordination complexes. The term“soluble redox species” excludes elemental metals and lone metal ions,especially those that are insoluble or sparingly soluble in water.

“Organotransition metal complex,” also referred to as “OTM complex,” isdefined as a complex where a transition metal is bonded to at least onecarbon atom through a sigma bond (formal charge of −1 on the carbon atomsigma bonded to the transition metal) or a pi bond (formal charge of 0on the carbon atoms pi bonded to the transition metal). For example,ferrocene is an OTM complex with two cyclopentadienyl (Cp) rings, eachbonded through its five carbon atoms to an iron center by two pi bondsand one sigma bond. Another example of an OTM complex is ferricyanide(Ill) and its reduced ferrocyanide (II) counterpart, where six cyanoligands (formal charge of −1 on each of the 6 ligands) are sigma bondedto an iron center through the carbon atoms of the cyano groups.

“Coordination complex” is defined as a complex having well-definedcoordination geometry, such as octahedral or square planar geometry.Unlike OTM complexes, which are defined by their bonding, coordinationcomplexes are defined by their geometry. Thus, coordination complexesmay be OTM complexes (such as the previously mentioned ferricyanide), orcomplexes where non-metal atoms other than carbon, such as heteroatomsincluding nitrogen, sulfur, oxygen, and phosphorous, are datively bondedto the transition metal center. For example, ruthenium hexaamine is acoordination complex having a well-defined octahedral geometry where sixNH₃ ligands (formal charge of 0 on each of the 6 ligands) are dativelybonded to the ruthenium center. A more complete discussion oforganotransition metal complexes, coordination complexes, and transitionmetal bonding may be found in Collman et al., Principles andApplications of Organotransition Metal Chemistry (1987) and Miessler &Tarr, Inorganic Chemistry (1991).

“Handheld device” is defined as a device that may be held in a humanhand and is portable. An example of a handheld device is the measurementdevice accompanying Ascensia® Elite Blood Glucose Monitoring System,available from Bayer HealthCare, LLC, Elkhart, Ind.

“On” is defined as “above” and is relative to the orientation beingdescribed. For example, if a first element is deposited over at least aportion of a second element, the first element is said to be “depositedon” the second. In another example, if a first element is present aboveat least a portion of a second element, the first element is said to be“on” the second. The use of the term “on” does not exclude the presenceof substances between the upper and lower elements being described. Forexample, a first element may have a coating over its top surface, yet asecond element over at least a portion of the first element and its topcoating can be described as “on” the first element. Thus, the use of theterm “on” may or may not mean that the two elements being related are inphysical contact with each other.

While various embodiments of the invention have been described, it willbe apparent to those of ordinary skill in the art that other embodimentsand implementations are possible within the scope of the invention.Accordingly, the invention is not to be restricted except in light ofthe attached claims and their equivalents.

The invention claimed is:
 1. An electrochemical biosensor fordetermining an analyte concentration in a fluid sample, the biosensorcomprising: a base; a first electrode on the base including a firstconductor and a first reagent layer, the first reagent layer including amediator having a deposition density of at most 5 μg/mm² and an enzymesystem having a deposition density of at most 0.5 μg/mm²; a secondelectrode on the base; and a lid, the lid and the base assisting informing an opening to receive the fluid sample, wherein the biosensorprovides a determined concentration value having a stability bias ofless than +/−10% after storage at 50° C. for 2 weeks when compared to acomparison biosensor stored at −20° C. for 2 weeks.
 2. The biosensor ofclaim 1, wherein the second electrode includes a second conductor and asecond reagent layer.
 3. The biosensor of claim 2, wherein the secondreagent layer is different in composition from the first reagent layer.4. The biosensor of claim 2, wherein the second reagent layer is thesame as the first reagent layer.
 5. The biosensor of claim 1, whereinthe average initial thickness of the first reagent layer is less than 8μm.
 6. The biosensor of claim 5, wherein the average initial thicknessof the first reagent layer is less than 4 μm.
 7. The biosensor of claim1, wherein the first reagent layer comprises at most 1.3 Units of theenzyme system.
 8. The biosensor of claim 1, wherein the stability biasis less than +/−5% after storage at 50° C. for 2 weeks when compared tothe comparison biosensor stored at −20° C. for 2 weeks.
 9. The biosensorof claim 1, wherein the first reagent layer further includes a polymericbinder selected from the group consisting of poly(ethylene oxide),polyvinyl alcohol, hydroxyethylene cellulose, carboxy methyl cellulose,and combinations thereof.
 10. The biosensor of claim 1, wherein themediator is selected from the group consisting of 3phenylimino-3H-phenothiazines, 3 phenylimino-3H phenoxazines, saltsthereof, acids thereof, derivatives thereof, and combinations thereof.11. An electrochemical biosensor for determining an analyteconcentration in a fluid sample, the biosensor comprising: a base; afirst electrode on the base including a first conductor and a reagentlayer, the reagent layer including a mediator having a depositiondensity of at most 2 μg/mm² and an enzyme system having a depositiondensity of at most 0.5 μg/mm²; a second electrode on the base; and alid, the lid and the base assisting in forming an opening to receive thefluid sample, wherein the biosensor provides a determined concentrationvalue having a stability bias of less than +/−10% after storage at 50°C. for 2 weeks when compared to a comparison biosensor stored at −20° C.for 2 weeks.
 12. An electrochemical biosensor for determining an analyteconcentration in a fluid sample, the biosensor comprising: a base; afirst electrode on the base including a first conductor and a reagentlayer, the reagent layer including a mediator having a depositiondensity of at most 5 μg/mm² and an enzyme system having a depositiondensity of at most 0.5 μg/mm²; a second electrode on the base; and alid, the lid and the base assisting in forming an opening to receive thefluid sample, wherein the biosensor provides a determined concentrationvalue having a hematocrit bias of less than +/−10% for whole bloodsamples including from 20 to 60% hematocrit.
 13. The biosensor of claim12, wherein the hematocrit bias is less than +/−5% for the whole bloodsamples including from 20 to 60% hematocrit.
 14. The biosensor of claim12, wherein the average initial thickness of the reagent layer is lessthan 8 μm.
 15. The biosensor of claim 12, wherein the reagent layerfurther includes a polymeric binder selected from the group consistingof poly(ethylene oxide), polyvinyl alcohol, hydroxyethylene cellulose,carboxy methyl cellulose, and combinations thereof.
 16. The biosensor ofclaim 12, wherein the mediator is selected from the group consisting of3 phenylimino-3H-phenothiazines, 3 phenylimino-3H phenoxazines, saltsthereof, acids thereof, derivatives thereof, and combinations thereof.17. An electrochemical biosensor for determining an analyteconcentration in a fluid sample, the biosensor comprising: a base; afirst electrode on the base including a first conductor and a reagentlayer, the reagent layer including a mediator having a depositiondensity of at most 2 μg/mm² and an enzyme system having a depositiondensity of at most 0.5 μg/mm²; a second electrode on the base; and alid, the lid and the base assisting in forming an opening to receive thefluid sample, wherein the biosensor provides a determined concentrationvalue having a hematocrit bias of less than +/−10% for whole bloodsamples including from 20 to 60% hematocrit.
 18. An electrochemicalbiosensor for determining an analyte concentration in a fluid sample,the biosensor comprising: a base; a first electrode on the base havingat least one first layer on a first conductor, the first layer includinga reagent layer comprising at most 5 μg/mm² of a mediator, and an enzymesystem at a deposition density of at most 0.5 μg/mm²; a second electrodeon the base; and a lid, the lid and the base assisting in forming anopening to receive the fluid sample, wherein the biosensor provides adetermined concentration value having an intercept to slope ratio of atmost 20 mg/dL.
 19. The biosensor of claim 18 wherein the intercept toslope ratio is at most 10 mg/dL.
 20. The biosensor of claim 19, whereinthe intercept to slope ratio is at most 6 mg/dL.
 21. The biosensor ofclaim 20, wherein the intercept to slope ratio is at most 1 mg/dL. 22.The biosensor of claim 18, wherein the average initial thickness of thereagent layer is less than 8 μm.
 23. The biosensor of claim 18, whereinthe reagent layer further includes a polymeric binder selected from thegroup consisting of poly(ethylene oxide), polyvinyl alcohol,hydroxyethylene cellulose, carboxy methyl cellulose, and combinationsthereof.
 24. The biosensor of claim 18, wherein the mediator is selectedfrom the group consisting of 3 phenylimino-3H-phenothiazines, 3phenylimino-3H phenoxazines, salts thereof, acids thereof, derivativesthereof, and combinations thereof.
 25. An electrochemical biosensor fordetermining an analyte concentration in a fluid sample, the biosensorcomprising: a base; a first electrode on the base having at least onefirst layer on a first conductor, the first layer including a reagentlayer comprising at most 2 μg/mm² of a mediator, and an enzyme system ata deposition density of at most 0.8 μg/mm²; a second electrode on thebase; and a lid, the lid and the base assisting in forming an opening toreceive the fluid sample, wherein the biosensor provides a determinedconcentration value having an intercept to slope ratio of at most 20mg/dL.
 26. A method of determining the concentration of an analyte in afluid sample, the method comprising: applying the fluid sample to anelectrochemical biosensor, the biosensor including a base, a firstelectrode, a second electrode and a lid, the first electrode on the baseincluding a first conductor and a reagent layer, the reagent layerincluding a mediator having a deposition density of at most 8 μg/mm² andan enzyme system having a deposition density of at most 0.8 μg/mm², thesecond electrode on the base, the lid and the base assisting in formingan opening to receive the fluid sample; applying a pulse sequence to thesample; determining current measurements from the sample responsive tothe pulse sequence; and determining the concentration of the analyte inthe sample from the current measurements, the determined concentrationhaving at least one of a stability bias of less than +/−10%, ahematocrit bias of less than +/−10% for whole blood samples over a 20 to60% hematocrit range, and an intercept to slope ratio of at most 20mg/dL.
 27. An electrochemical biosensor for determining an analyteconcentration in a fluid sample, the biosensor comprising: a base; afirst electrode on the base including a first conductor and a reagentlayer, the reagent layer including a reagent layer comprising from 1.72to 2 μg/mm² of a mediator and an enzyme system having a depositiondensity from 0.42 to 0.8 μg/mm²; a second electrode; and a lid, the lidand the base assisting in forming an opening to receive the fluidsample.
 28. An electrochemical biosensor for determining an analyteconcentration in a fluid sample, the biosensor comprising: a base; afirst electrode on the base including a first conductor and a reagentlayer, the reagent layer including a reagent layer comprising from 1.72to 8 μg/mm² of a mediator and an enzyme system having a depositiondensity from 0.42 to 0.8 μg/mm²; a second electrode; and a lid, the lidand the base assisting in forming an opening to receive the fluidsample.